Cardiac support device

ABSTRACT

A highly compliant and elastic cardiac support device is provided. The device is constructed from a biocompatible material is applied to an external surface of a heart. The device can be used to resist dilatation of the heart, to provide acute wall support, or to enhance reduction in the size of the heart using stored potential energy, without interfering with systolic contraction.

CROSS REFERENCE TO RELATED APPLICATIONS

This application is a continuation-in-part of U.S. patent applicationSer. No. 09/593,251, filed Jun. 13, 2000, now U.S. Pat. No. 6,482,146.

BACKGROUND OF THE INVENTION

Congestive heart disease is a progressive and debilitating illnesscharacterized by a progressive enlargement of the heart. As the heartenlarges, the heart must perform an increasing amount of work to pumpblood each heartbeat needed for metabolism. In time, the heart becomesso enlarged the heart cannot adequately supply blood. An afflictedpatient is fatigued, unable to perform even simple exerting tasks andexperiences pain and discomfort. Further, as the heart enlarges, theinternal heart valves may not adequately close. This may impair thefunction of the valves and further reduce the heart's ability to supplyblood. Importantly, there is no cure for congestive heart disease.

Congestive heart failure has an enormous societal impact. In the UnitedStates alone, about five million people suffer from the disease.Alarmingly, congestive heart failure is one of the most rapidlyaccelerating diseases (about 400,000 new patients in the United Stateseach year). Economic costs of the disease have been estimated at $38billion annually. Not surprising, substantial effort has been made tofind treatments for congestive heart disease.

Cardiomyoplasty is one potential treatment for moderate stage congestiveheart disease. In this procedure, the latissimus dorsi muscle (takenfrom the patient's shoulder) is wrapped around the heart and chronicallypaced to assist contraction of the heart during systole.

One study speculates that an artificial elastic “sock” could be used inplace of the latissimus dorsi in adynamic cardiomyoplasty. Kass et al.,“Reverse Remodeling from Cardiomyoplasty in Human Heart Failure,”Circulation 91:9, May 1, 1995. Another study demonstrates that BardMarlex sheets can be used to wrap the heart as a substitute to thelatissimus dorsi in adynamic cardiomyoplasty. Oh et al., “The Effects ofProsthetic Cardiac Binding and Adynamic Cardiomyoplasty in a Model ofDilated Cardiomyoplasty,” Journal of Thoracic Cardiovascular Surgery,116:1, July 1998. German Utility Model Patent Application DE 295 17 393U1 describes a pericardium prosthesis made from a biocompatible,non-expansible material, or at least hardly expansible material thatsurrounds the heart to prevent over-expansion of the heart wall. PCTapplication WO 98/58598 describes an elastic pouch for at leastpartially enveloping a heart. Commonly assigned U.S. Pat. No. 5,702,343to Alferness dated Dec. 30, 1997 teaches a jacket to constrain cardiacexpansion during diastole. Other teachings include those of commonlyassigned U.S. Pat. No. 6,123,662 and those of U.S. Pat. ApplicationPublication No. US 2002/0019580.

SUMMARY OF THE INVENTION

The invention provides a device for treating cardiac disease. Accordingto the invention, a highly compliant and elastic device, constructedfrom a biocompatible material is applied to an external surface of aheart. The device can be used to resist dilatation of the heart, provideacute wall support, and/or to enhance reduction in the size of the heartusing stored potential energy, without interfering with systoliccontraction. According to the invention, the device has a compliancethat is greater than a compliance of a normal myocardium, morepreferably, the device has a compliance greater than a compliance of anormal latissimus dorsi. Considering that stiffness is the inverse ofcompliance, the uniaxial stiffness of the material is generally lessthan about 0.5 lbs/in (i.e. load per width of device) when subject to auniaxial load at a strain of less than 30%, more typically between about0.05 lbs/in and about 0.2 lbs/in. An alternative way to examinecompliance for a device that is applied to an enclosed volume is basedon a 3-dimensional volumetric compliance. The 3-dimensional complianceof the device typically allows at least a 3% increase in volume forevery 1 mm Hg change in applied device pressure. More typically, thematerial of the device has a 3-dimensional volumetric compliance betweenabout 5%/mm Hg and about 15%/mm Hg. The device typically has an elasticrecovery of at least about 50%, but 70% to 100% is preferable.

In one embodiment, the material of the device is sized to be smallerthan the external surface of the heart to which it is applied. Inanother embodiment, the material of the device is sized to be largerthan the external surface of the heart to which it is applied andadapted to be sized by adjustment during implantation. The material canbe configured as a jacket for covering both ventricles, one ventricle,ventricles and atria, atria or as a patch covering a portion of achamber.

The invention also provides a method for treating a cardiac disease. Themethod includes a step of surgically accessing the heart; and applying acardiac support device to an external surface of the heart. The methodcan be used to resist dilatation of the heart, acutely support the wallof the heart, and/or to enhance reduction in a size of the heart usingstored potential energy.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a schematic cross-sectional view of a normal, healthy humanheart shown during systole;

FIG. 1A is the view of FIG. 1 showing the heart during diastole;

FIG. 1B is a view of a left ventricles of a healthy heart as viewed froma septum and showing a mitral valve;

FIG. 2 is a schematic cross-sectional view of a diseased human heartshown during systole;

FIG. 2A is the view of FIG. 2 showing the heart during diastole;

FIG. 2B is the view of FIG. 1B showing a diseased heart;

FIG. 3 is a schematic showing the theory of operation of a cardiacsupport device.

FIG. 4 is a perspective view of an embodiment of a cardiac supportdevice according to the present invention;

FIG. 4A is a side elevation view of a diseased heart in diastole withthe device of FIG. 4 in place;

FIG. 5 is a perspective view of another embodiment of a cardiac supportdevice with the apex open according to the present invention;

FIG. 5A is a side elevation view of a diseased heart in diastole withthe device of FIG. 5 in place;

FIG. 6 is a plan view of an alternate embodiment of a cardiac supportdevice;

FIG. 7 is a side elevation view of a diseased heart in diastole withanother embodiment of a cardiac support device in place;

FIG. 8 is a cross-sectional view of a device of the present inventionoverlying a myocardium and with the material of the device gathered fora snug fit;

FIG. 9 is an enlarged simplified view of a the fabric of a knitconstruction at rest, suitable for use in the device of this invention;

FIG. 10 shows compliance curves (pressure versus % volume change) for alower compliance (A) and a higher compliance (B) device;

FIG. 11 shows compliance curves (pressure versus % volume change) for alower compliance device (A) and a higher compliance device (B) atimplant;

FIG. 12 is a device elastic potential energy plot comparing the workenergy stored at implant for a less compliant and elastic device (A) anda highly compliant and elastic device (B);

FIG. 13 is a plot of the change in left ventricle end diastolicdimension (LVEDD) over time from clinical studies;

FIG. 14 is a plot of device loading (♦) and unloading (▪) fordetermining elastic recovery; and

FIG. 15 is a schematic of the mechanical roles in cardiac support devicetherapy;

FIG. 16 is a plot of LVEDV (left ventricular end diastolic volume)change results from a pre-clinical animal study with a horizontal axisdepicting time and with a vertical axis depicting change in LVEDV frompre-implant conditions.

DESCRIPTION OF THE PREFERRED EMBODIMENT

A. Congestive Heart Disease

With initial reference to FIGS. 1 and 1A, a normal, healthy human heartH′ is schematically shown in cross-section and will now be described inorder to facilitate an understanding of the present invention. In FIG.1, the heart H′ is shown during systole (i.e., high left ventricularpressure during the ejection phase). In FIG. 1A, the heart H′ is shownduring diastole (i.e., low left ventricular pressure during therelaxation phase).

The heart H′ is a muscle having an outer wall or myocardium MYO′ and aninternal wall or septum S′. The myocardium MYO′, septum S′ and valveplane VP′ define four internal heart chambers, including a right atriumRA′, a left atrium LA′, a right ventricle RV′ and a left ventricle LV′.The heart H′ has a length measured along a longitudinal axis AA′-BB′from an upper end or base B′ to a lower end or apex A′.

The heart H′ can be visualized as having an upper portion UP′ and alower portion LP′, separated by the valve plane VP′. On the externalsurface of the heart, the upper portion UP′ and lower portion LP′ meetat a circumferential groove commonly referred to as the A-V groove AVG′right and left atria RA′, LA′ reside in an upper portion UP′ of theheart H′ adjacent the base B′. The right and left ventricles RV′, LV′reside in a lower portion LP′ of the heart H′ adjacent the apex A′. Theventricles RV′, LV′ terminate at ventricular lower extremities LE′adjacent the apex A′ and spaced therefrom by the thickness of themyocardium MYO′.

Extending away from the upper portion UP′ are a plurality of major bloodvessels communicating with the chambers RA′, RV′, LA′, LV′. For ease ofillustration, only the superior vena cava SVC′ and a left pulmonary veinLPV′ are shown as being representative.

The heart H′ contains valves to regulate blood flow between the chambersRA′, RV′, LA′, LV′ and between the chambers and the major vessels (e.g.,the superior vena cava SVC′ and a left pulmonary vein LPV′). For ease ofillustration, not all of such valves are shown. Instead, only thetricuspid valve TV′ between the right atrium RA′ and right ventricle RV′and the mitral valve MV′ between the left atrium LA′ and left ventricleLV′ are shown as being representative.

The valves are secured, in part, to the myocardium MYO′ in a region ofthe A-V groove AVG′ and referred to as the valve plane VP′ or valvularannulus VA′. The valves TV′ and MV′ open and close through the beatingcycle of the heart H′.

FIGS. 1 and 1A show a normal, healthy heart H′ during systole anddiastole, respectively. During systole (FIG. 1), the myocardium MYO′ iscontracting and the heart assumes a shape including a generally conicallower portion LP′. During diastole (FIG. 1A), the heart H′ is expandingand the conical shape of the lower portion LP′ bulges radially outwardly(relative to axis AA′-BB′).

The motion of the heart H′ and the variation in the shape of the heartH′ during contraction and expansion is complex. The amount of motionvaries considerably throughout the heart H′, although the externaldimension of the heart H′ generally reduces from about 4% to about 10%from end diastole to end systole. The motion includes a component whichis parallel to the axis AA′-BB′ (conveniently referred to aslongitudinal expansion or contraction). The motion also includes acomponent perpendicular to the axis AA′-BB′ (conveniently referred to ascircumferential expansion or contraction).

Having described a healthy heart H′ during systole (FIG. 1) and diastole(FIG. 1A), comparison can now be made with a heart deformed bycongestive heart disease. Such a heart H is shown in systole in FIG. 2and in diastole in FIG. 2A. All elements of diseased heart H are labeledidentically with similar elements of healthy heart H′ except only forthe omission of the apostrophe in order to distinguish diseased heart Hfrom healthy heart H′.

Comparing FIGS. 1 and 2 (showing hearts H′ and H during systole), thelower portion LP of the diseased heart H has lost the tapered conicalshape of the lower portion LP′ of the healthy heart H′. Instead, thelower portion LP of the diseased heart H bulges outwardly between theapex A and the A-V groove AVG. So deformed, the diseased heart H duringsystole (FIG. 2) resembles the healthy heart H′ during diastole (FIG.1A). During diastole (FIG. 2A), the deformation is even more extreme.

As a diseased heart H enlarges from the representation of FIGS. 1 and 1Ato that of FIGS. 2 and 2A, the heart H becomes a progressivelyinefficient pump. Therefore, the heart H requires more energy to pumpthe same amount of blood. Continued progression of the disease resultsin the heart H being unable to supply adequate blood to the patient'sbody and the patient becomes symptomatic insufficiency. In contrast to ahealthy heart H′, the external dimension of the diseased heart Hgenerally reduces from about 4% to about 6% from end diastole to endsystole.

For ease of illustration, the progression of congestive heart diseasehas been illustrated and described with reference to a progressiveenlargement of the lower portion LP of the heart H. While suchenlargement of the lower portion LP is most common and troublesome,enlargement of the upper portion UP may also occur.

In addition to cardiac insufficiency, the enlargement of the heart H canlead to valvular disorders. As the circumference of the valvular annulusVA increases, the leaflets of the valves TV and MV may spread apart.After a certain amount of enlargement, the spreading may be so severethe leaflets cannot completely close (as illustrated by the mitral valveMV in FIG. 2A). Incomplete closure results in valvular regurgitationcontributing to an additional degradation in cardiac performance. Whilecircumferential enlargement of the valvular annulus VA may contribute tovalvular dysfunction as described, the separation of the valve leafletsis most commonly attributed to deformation of the geometry of the heartH. This is best described with reference to FIGS. 1B and 2B.

FIGS. 1B and 2B show a healthy and diseased heart, respectively, leftventricle LV′, LV during systole as viewed from the septum (not shown inFIGS. 1B and 2B). In a healthy heart H′, the leaflets MVL′ of the mitralvalve MV′ are urged closed by left ventricular pressure. The papillarymuscles PM′, PM are connected to the heart wall MYO′, MYO, near thelower ventricular extremities LE′, LE. The papillary muscles PM′, PMpull on the leaflets MVL′, MVL via connecting chordae tendineae CT′, CT.Pull of the leaflets by the papillary muscles functions to prevent valveleakage in the normal heart by holding the valve leaflets in a closedposition during systole. In the significantly diseased heart H, theleaflets of the mitral valve may not close sufficiently to preventregurgitation of blood from the ventricle LV to the atrium duringsystole.

As shown in FIG. 1B, the geometry of the healthy heart H′ is such thatthe myocardium MYO′, papillary muscles PM′ and chordae tendineae CT′cooperate to permit the mitral valve MV′ to fully close. However, whenthe myocardium MYO bulges outwardly in the diseased heart H (FIG. 2B),the bulging results in displacement of the papillary muscles PM. Thisdisplacement acts to pull the leaflets MVL to a displaced position suchthat the mitral valve cannot fully close.

B. Cardiac Support Therapy

In general, cardiac support therapy uses a “passive” mechanical implantto support the heart and resist circumferential expansion of the heartduring diastole and without actively assisting contraction duringsystole. Herein, the term “passive” is used to contrast the device withan “active assist” device which uses supplied energy in order tooperate, such as devices that assist the heart in pumping blood flowinto the aorta, for example, left ventricular assist devices (“LVAD”)and total artificial hearts (“TAH”). However, the device of theinvention does have some mechanical components that involve energy inputinto the system, and therefore are not entirely “passive.” As usedherein, the term “active” refers to a device wherein energy is added tothe system on an ongoing basis. In contrast, a “passive” device, as usedherein, may use stored or potential energy. The potential energy storedin the device is generally attributable to energy that is input when thedevice is fit on the heart, which in part is due to mechanicalproperties of the device (such as compliance and elasticity). The devicecan be thought of as having stored energy, similar to a pre-loadedspring. However, in contrast to an “active” device, once the device ofthe invention is implanted, no additional energy is added continually.The mechanical components of the device that involve energy input aredescribed in detail below.

It is believed that the cardiac support device stimulates aphysiological response due to a mechanical effect, a tissue-materialinteraction, or some combination thereof. While the physiologicalresponse can be difficult to predict, the mechanical interactions aremore straightforward. FIG. 3 is a schematic showing how a cardiacsupport device interrupts the cycle of heart failure by disruptingexcessive ventricular dilatation (i.e., abnormal dilation) duringdiastolic filling. Briefly, following an injury to the myocardium, theheart's function may be reduced (A). This stimulates a compensatoryresponse of ventricular dilatation (B) to improve output. However,ventricular dilatation causes increased wall stress and stretch (D),which then triggers neurohormonal activation (C), leading to modifiedgene expression (E) which in turn leads to structural and functionalchanges in the myocardium. These changes are also referred to asventricular remodeling (F). This further reduces cardiac functioncausing the cycle to repeat with additional compensatory dilatation.FIG. 3 illustrates the potential benefits of a cardiac support deviceproviding wall support and resistance to ventricular dilatation. Acardiac support device reduces the myocardial wall stress and stretch(H), which helps to break the heart failure cycle and leads to improvedefficiency (I), reverse remodeling (J) and ultimately improved cardiacfunction (K).

The mechanical effects that help interrupt the ongoing ventriculardilatation in the heart failure cycle can be divided into at least threemechanisms or components: (1) dilatation constraint (acute and/orchronic); (2) acute wall support; and (3) chronic potential energyrelease. As used herein, “dilatation constraint” means resistingexpansion or dilatation of a heart that would result in a damagingincrease in the volume of the heart. As used herein, “acute wallsupport” means reducing stress on the wall of the heart or supportingthe internal pressure (i.e. reducing transmural wall pressure) of theheart by off-loading the heart acutely, at the time of the deviceimplant. As used herein, “chronic potential energy release” refers tothe potential energy of the device that is available (and released) toencourage reduction in the size of the heart, in terms of volume and/ordimension over time following implant of the device. Many of the deviceparameters, whether used for dilatation constraint, acute wall supportand/or potential energy release, overlap. However, for the sake ofclarity the design characteristics and/or the method for implanting adevice for each of the three mechanisms or components will be discussedseparately below. FIG. 15 is a schematic diagram showing the mechanicalmodes for the device mechanisms and how they tie into the biologicalresponses to support the theory of operation in FIG. 3.

1. Dilatation Constraint (Acute and/or Chronic)

Dilatation constraint refers to the resistance the device provides toshort-term transient dilatation and/or chronic cardiac dilatation, inparticular excessive ventricular dilatation. Generally, dilatationconstraint does not require energy input into the device. The device istypically adjusted to conform to the epicardial surface, to resistfurther dilatation of the heart. “Acute” dilatation constraint refers toresistance to cardiac dilatation from short term loading such asexercise loading. Exercise loading refers to the loading that occurswhen a person performs a physical activity such as exercise. In exerciseloading, the heart increases it's volume to provide more output usingthe Frank-Starling relationship. However, the increased volume resultsin increased end diastolic loading of the ventricular wall due to theLaw of LaPlace. The Law of LaPlace is based on the concept that thelarger the vessel radius, the larger the wall tension required towithstand a given internal fluid pressure. Larger ventricular chambervolumes generally correspond to larger chamber radii. “Chronic”dilatation constraint refers to resistance to continued dilatation dueto prolonged volume loading and cardiac remodeling. Increased volumeloading can also result from the intake of fluids, which is notdiscussed here, or kidney damage that is also associated with heartdisease.

During dilatation constraint, the device reduces the ventricular wallstress and stretch increase that accompany acute and continueddilatation in heart failure. Both the reduction in ventricular wallstress and stretch increase are “myocardial displacement dependent.” Asused herein, the phrase “myocardial displacement dependent” means thatthe amount of support or loading provided by the cardiac support deviceis dependent on the amount of myocardial wall dimensional dilatationcaused by disease progression or excessive loading. For this mechanicalcomponent, the compliance of the device can be an importantcharacteristic. Generally, the compliance of the device can be importantfor acute loading before the fibrosis encapsulates the device and forlong-term chronic dilatation. However, device compliance tends to beless important for acute loading after the fibrosis develops. Generally,lower compliance (i.e. higher stiffness) tends to provide moreresistance and support for dilatation.

Generally, support devices such as those mentioned in the Backgroundsection of this application have focused on the mechanism of dilatationconstraint.

Generally, when the device is used for dilatation constraint, the device10 surrounds the myocardium MYO, as shown, for example, in FIG. 4. Asused herein, “surround” means that the device provides reduced expansionof the heart wall during diastole by applying constraining surfaces atleast at diametrically opposing aspects of the heart. Generally, thediametrically opposed surfaces are interconnected, for example, by acontinuous material that can substantially encircle the external surfaceof the heart.

In one embodiment, the device is configured as a jacket 10 that definesa volume 16. Preferably the volume 16 is substantially the same size asor larger than the volume of the heart H, in particular the lowerportion LP of the heart, at the completion of systolic contraction suchthat the jacket 10 exerts no or only a slight pressure on the heart atend systole. Generally, the jacket 10 is adjusted such that the jacket10 resists enlargement of the heart H during diastole withoutsignificantly assisting contraction during systole. At time ofplacement, the device preferably exerts no or only a small pressure onthe heart H at end diastole of less than 10 mm Hg, more preferably lessthan or equal to 5 mm Hg, most preferably less than or equal to 2 mm Hg.Such pressure may be determined by comparing load to the rightventricular end diastolic pressure.

To permit the jacket 10 to be easily placed on the heart H, the volumeand shape of the jacket 10 may be larger than the lower portion LPduring diastole. So sized, the jacket 10 may be easily slipped aroundthe heart H. Once placed, the jacket's volume and shape can be adjustedfor the jacket 10 to snugly conform to the external geometry of theheart H during diastole. For example, excess material of the jacket 10can be gathered and sutured S″ (FIG. 8) to reduce the volume of thejacket 10 and conform the jacket 10 to the shape of the heart H duringdiastole. This shape represents an adjusted volume. The jacket 10resists enlargement of the heart H beyond the adjusted volume withoutinterfering with contraction of the heart H during systole. As analternative to the gathering shown in FIG. 8, the jacket 10 can beprovided with other ways of adjusting volume. For example, as disclosedin U.S. Pat. No. 5,702,343, the jacket can be provided with a slot. Theedges of the slot can be drawn together to reduce the volume of thejacket. The volume of the jacket can be adjusted prior to, during, orafter application of the device to the heart.

Although, for dilatation constraint, the device is generally adjusted toa snug fit as described above, it is also possible to obtain thebenefits of dilatation constraint using a device that defines a volumethat is smaller than the volume of the portion of the heart H on whichit is to be placed at end diastole. In this embodiment, the device isstretched in order to place it around the heart H, such that thecompliance of the jacket 10 material and the amount of expansion of thematerial at end diastole determine the fit of the device without anyfurther adjustment.

2. Acute Wall Support

Acute wall support refers to a more immediate effect of the cardiacsupport device. Generally, acute wall support is obtained by adjustingthe device such that the device applies an external pressure to theheart. If desired, the device can be adjusted to provide a dimensionalreduction in the heart size. For example, the device may be adjusted toslightly reduce cardiac dimension at the time of implantation,preferably, no more than 10% reduction in internal Left Ventricular EndDiastolic Dimension (LVEDD). Thus, rather than just reducing theincrease in wall stress and stretch due to dilatation constraint, energyis actively input at the time of implantation to reduce the load on thewall acutely. However, after the device is placed on the heart, nofurther external energy is added. Thus, the device is still considered a“passive” device. Acute wall support reduces wall stress (loaddependent) and reduces wall stretch (myocardial displacement dependent).As used herein, the phrase “load dependent” means that the reduction inwall stress is dependent on the amount of load applied, independent ofthe amount of dimensional change. The reduction in end diastolic wallstress is based on the change in transmural heart wall pressure. Incontrast, the amount of reduced wall stretch is related to thedimensional reduction in the heart size. In contrast to a dilatationconstraint mechanism, energy is input at the time of implantation foracute wall support and the material compliance is less important.However, as mentioned previously, these components overlap such that thebenefits from both dilatation constraint and acute wall support can berealized from the same device.

If the device is configured as a jacket 10, it may be desirable to havea volume and shape that is larger than the lower portion LP duringdiastole so that the jacket 10 may be easily slipped around the heart Hand adjusted (as with dilatation constraint). However, it may also bedesirable to use a device with a volume and shape that is smaller thanthe lower portion LP of the heart H during diastole. In this embodiment,the compliance of the jacket 10 and expansion at diastole determine thefit, without additional adjustment. When selecting or adjusting thejacket 10 for acute wall support, care should be taken to avoidimpairing normal cardiac function. During diastole, the left ventricleLV fills with blood. If the jacket 10 is too tight, the left ventricleLV may not adequately expand and left ventricular filling pressure mayrise. Furthermore, if the device encloses both ventricles such as inFIG. 4A, care should be taken when selecting or adjusting the jacket 10,because the wall of the right ventricle RV tends to be thinner than thewall of the left ventricle LV and the pressure in the right ventricle RVtends to be lower than the pressure in the left ventricle LV. Preferablythe pressure exerted by the jacket 10 on the heart H at end diastole isnot greater than the end diastolic pressure of the right ventricle RV.If the pressure exerted by the jacket 10 is greater than the pressure ofthe right ventricle RV, expansion and/or filling of the right ventricleRV may be compromised. However, for a device that is applied to only oneof the ventricular chambers such as the Left Ventricle LV as shown inFIG. 7, the pressure exerted by the jacket 10 at end diastole ispreferably less than the end diastolic pressure of the LV.

Generally a jacket 10 that imposes between about a 5% to about a 10%reduction in LVEDD (left ventricle end diastolic dimension) serves toreduce cardiac volume without compromising cardiac function. Preferably,the jacket 10 exerts pressure at end diastole between about 2 mm Hg andabout 20 mm Hg, more preferably between about 5 mm Hg and about 15 mmHg, and most preferably between about 5 mm Hg and about 10 mm Hg,depending on the internal end diastolic pressures of the heart chambers.The jacket may be designed with multiple sections with differentcompliances and pressures for a specific heart chamber.

3. Chronic Potential Energy Release

In addition to dilatation constraint and acute wall support, the cardiacsupport device may also be able to use stored potential energy toenhance heart size reduction over time. Based on the Law of LaPlace,reduced heart size will reduce the myocardial wall load for a giveninternal chamber pressure. The chronic potential energy releasemechanism and the device properties that enable the size reduction arekey aspects of this invention. Generally, the potential energy of thedevice is due to the fabric being stretched at the time of implantation.Typically, the device is selected and/or adjusted (if necessary) to havea “resting” size and/or volume that are smaller than that of theenlarged heart to which it is applied. Preferably the “resting” size ofthe device is approximately the same size as the heart in a healthystate or some other desired target size. As used herein the term“resting” means that the fibers of the fabric are in a relaxed statesuch that energy is not required to keep the fibers in the “resting” or“relaxed” state. When the material is “stretched” to accommodate theenlarged heart, work energy must be input to create the “stretched”configuration. The amount of energy input and stored in the device isbased on the amount of strain (or stretch) and the load required toobtain that strain. According to one aspect of the invention, thematerial is stretched during implantation, wherein the stretchingprovides the material with potential energy that can be used to enhancereduction in the size of the heart. In one embodiment, the material isstretched to provide a stretched volume that is at least 20% greaterthan the resting volume, more preferably, the material is stretched toprovide a stretched volume that is about 40% greater than the restingvolume, more preferably about 60% greater than the resting volume. Themaximum stretch should be based on the limit of heart volume reductiondesired. Similarly, a patch device that covers a small area (i.e. FIG.6) rather than encapsulating a volume may have similar stretch targetsbased on area or length rather than volume.

Again, as with acute wall support, care should be taken to avoidexerting too much pressure on the heart, such that cardiac function isimpaired. For this mechanical mechanism or component, the devicepreferably exerts pressures similar to those described for the acutewall support mechanism. Lower pressures may be effective and morepreferred depending on the compliance and elasticity of the device andthe desired level of stored potential energy.

Both the compliance and elasticity of the material are importantparameters for the chronic potential energy release mechanism.Compliance refers to the ability of the device to deform under load. Inengineering terms it is the inverse of stiffness. Elasticity refers tothe ability of the device to return to its original dimension uponunloading after being deformed by a load. The compliance of the deviceand the load applied determine the amount of energy added to the systemat the time of implant. However, the device elasticity determines thenew resting state of the device after it has been stretched out forimplantation, and how much stored potential energy can be released fromunloading the device. Once the device is implanted, the heart willgenerally reduce in size over time to reduce the external load appliedby the device. The amount of potential energy stored and the elasticityin the device will affect how much the device can mechanically reduceand reshape the heart from a dilated size to possibly a normal size. Adevice having high compliance and high elasticity is generally preferredfor this mechanism to increase the amount of potential energy that canbe stored and recaptured.

It is noted that the tissue response to the implanted device may causethe device to be encapsulated in a thin layer of fibrosis. Thecollagenous fibrotic tissue can be remodeled when it is subjected tochronic loads. Thus, after the device is encapsulated by fibrosis, thecomposite compliance of the fibrosis and device may be reduced forshort-term transient loads. However, for long-term loads such asreduction in the heart size due to the chronic potential energy releaseof the device, fibrosis is believed to have only a minor orinsignificant impact on the compliance and elasticity of the device.Over time the fibrosis is unlikely able to support the load from theheart or the device. Thus, the fibrosis tends to remodel as the heartreduces over time and the compliance and elasticity of the devicecontinue to drive the mechanical reduction in heart size until thepotential energy of the device is fully released or the heart sizestabilizes.

The chronic potential energy release mechanism of the device reduceswall stress and reduces wall stretch, both of which are myocardialdisplacement dependent. As the device mechanically causes the heart toreduce in size, the heart wall stress and stretch reduce due to thechange in geometry. The chronic potential energy mechanism wasdemonstrated in a pre-clinical animal model. FIG. 16 with the earlyresults of an animal study using canines with failing hearts showssignificantly larger left ventricular end diastolic volume (LVEDV)reduction in two animals implanted with a higher compliance device (A)when compared to six animals implanted with the current lower compliantdevice (B). All animals were implanted with similar loading and littleto no acute reduction at the time of implant. The additional potentialenergy stored in the high compliance devices was able drive the sizereductions by over 3 times more volume.

C. Cardiac Support Device.

The invention provides a device having a compliance and/or elasticity torender it suitable for use for one or more of the following treatments:resisting enlargement of the cardiac dimension (dilatation constraint),offloading stress from the myocardial wall (acute wall support), andenhancing reduction in cardiac dimension (chronic potential energyrelease).

Generally, the device is configured to cover at least part of theepicardial surface, typically at least the ventricles. As used herein,the term “cover” means that the device is in contact with an externalsurface and applies a force on the surface of the heart. Generally, thedevice contacts an epicardial surface of the heart, but it can also beapplied over the pericardium.

A device that “covers” the lower extremities of the heart may beconstructed as a continuous material that can substantially encircle, or“surround”, the external surface of the lower extremities of the heart(See, FIGS. 4, 4A, 5, 5A). In an alternate embodiment, the deviceprovides for localized support of the heart, particularly duringdiastole. According to this embodiment, a device 10 may be configured asa “patch” (See, FIG. 6). A patch may be useful to provide dilatationconstraint or acute wall support over a localized area of injury such asan acute myocardial infarction (AMI) or a wall aneurysm. In the case ofan aneurysm, it may be advantageous to take advantage of the chronicpotential energy release mechanism to restore the wall shape over time.When discussing a “patch”, the size of the patch is selected to cover anarea of the epicardial surface of the heart without completelysurrounding the circumference of the heart. In yet another embodiment,the device may be configured to cover only a left or right ventricle(See, FIG. 7). Typically, in this embodiment, the device is attached tothe heart proximate the septal wall S′. If desired, the device can beconstructed from material having one or more compliances or beconstructed as one or more separate components. The mechanicalcharacteristics of each component may be designed to specifically targetone or more of the mechanical mechanisms of device therapy previouslydescribed. With reference now to FIGS. 4, 4A, 5 and 5A, the device ofthe present invention is shown as a jacket 10 of flexible, biologicallycompatible material. As used herein, the term “biologically compatiblematerial” refers to material that is biologically inert such that thematerial does not adversely result in excessive injurious responses suchas chronic inflammation which would adversely affect the myocardium andpotentially surrounding tissues.

A jacket 10 is an enclosed material having upper and lower ends 12, 14.The jacket 10, 10′ defines an internal volume 16, 16′ which iscompletely enclosed but for the open ends 12, 12′ and 14′. In theembodiment of FIG. 4, lower end 14 is closed. In the embodiment of FIG.5, lower end 14′ is open. In both embodiments, upper ends 12, 12′ areopen. Throughout this description, the embodiment of FIG. 4 will bediscussed. Elements in common between the embodiments of FIGS. 4 and 5are numbered identically with the addition of an apostrophe todistinguish the second embodiment and such elements need not beseparately discussed.

The jacket 10 is dimensioned with respect to a heart H to be treated.Specifically, the jacket 10 is sized for the heart H to be enclosedwithin the volume 16. The jacket 10 can be slipped around the heart H.The jacket 10 has a length L between the upper and lower ends 12, 14sufficient for the jacket 10 to enclose the lower portion LP. In oneembodiment, the upper end 12 of the jacket 10 extends at least to thevalvular annulus VA and further extends to the lower portion LP toenclose at least the lower ventricular extremities LE. If desired, thejacket 10 may be sized so that the upper end 12 resides in the A-Vgroove AVG. Where it is desired to treat the upper portion UP, thejacket 10 may be extended to cover the upper portion UP.

After the jacket 10 is positioned on the heart H as described above, thejacket 10 is secured to the heart. Preferably, the jacket 10 is securedto the heart H through sutures. The jacket 10 is sutured to the heart Hat suture locations 15 circumferentially spaced along the upper end 12.While a surgeon may elect to add additional suture locations to preventshifting of the jacket 10 after placement, the number of such locations15 is preferably limited so that the jacket 10 does not restrictcontraction of the heart H during systole. Other attachment methods suchas staples or clips may be acceptable as an alternative to sutures alongthe upper end 12.

The jacket 10 can be adjusted to provide the appropriate fit afterplacement around the heart. Alternatively, the jacket 10 can be sized toobtain the appropriate fit based on device compliance, desired level offit and the size of the portion of the heart H the device is intended tocover.

a. Compliance

As used herein, the term “compliance” refers to the load required todeform the material. As mentioned previously, in the field ofengineering it is the inverse of stiffness. The compliance can bedescribed in terms of displacement (inches or centimeters), strain(inch/inch or cm/cm) or volume (in³, cm³ or ml) per a unit load (poundsor kilograms) or pressure (psi or mm Hg). The compliance of the cardiacsupport device can have a significant impact on the mechanical mechanismand effectiveness in the therapy, as well as allowing it to be stretchedfor accommodating an enlarged heart. It should also be noted thatcompliance is not necessarily constant over a given range ofdisplacement. In fact, compliance that decreases with increased stretchis a common characteristic of many materials.

Due to the Frank-Starling behavior of the heart, a cardiac supportdevice that has less compliance than the myocardium at smalldeformations may not be desirable. Generally, the cardiac output demandfor the heart changes depending on physical activity. To increase thecardiac output according to the Frank-Starling mechanism, the pre-loador end diastolic volume of the heart is increased such that the musclefibers are temporarily stretched. The stretching of the muscle fibershelps increase heart capacity and myocyte contractility and thereforecardiac output. If a cardiac support device with less compliance thanthe myocardium is applied to the surface of the heart, the heart may notbe able to utilize the Frank-Starling mechanism effectively withoutincreasing ventricular filling pressure. Thus, ventricular filling maybe negatively impacted and mimic a cardiovascular disease known asconstrictive pathologies. Thus, a cardiac support device with a highercompliance than the myocardium is generally preferred.

Other evidence indicating that cardiac support with higher compliancemay be preferable can be found in examining the compliance of the normalpericardium and the latissimus dorsi muscle used to wrap the heart forcardiomyoplasty. The stiffness of living myocardium and latissimus dorsimuscle is complex and has both active and passive elements. Forsimplicity, only the passive elements will be examined.

Table 1 contains a comparison of passive stiffness of myocardial tissue,pericardial tissue, latissimus dorsi muscle tissue and a sample cardiacsupport device that has been described in previous patents (i.e. U.S.Pat. No. 6,085,754 and International patent application publication No.PCT WO 01/95830) (this sample cardiac support device is referred toherein as the “prior knit device”). The values in Table 1 were derivedbased on uniaxial loads at strains less than 30%. As shown in Table 1,the pericardial tissue is much more compliant than the myocardium forlow strains, but stiffens at higher strains to become less compliantthan the myocardium. If the pericardium was as stiff for low strains ashigher strains, ventricular filling would probably be impaired, similarto constrictive pericarditis or cardiac tamponade. The stiffnesscomparisons provided in Table 1 illustrate that the latissimus dorsi(sometimes used to wrap the heart for cardiomyoplasty) is also morecompliant than the myocardial tissue. The sample cardiac support device(i.e., the prior knit device) also has a greater compliance than themyocardium and similar, but slightly less compliance than the latissimusdorsi muscle. As mentioned earlier in the Background section, Oh et al.used a very non-expansible material to wrap around the heart known asBard Marlex. Although the Bard Marlex helped to limit progressivedilatation in this study, it was not as effective as the latissimusdorsi muscle in adynamic cardiomyoplasty. The uniaxial stiffness of BardMarlex has been measured to be less compliant than the myocardium asshown in Table 1. The data in Table 1 thus supports the concept that amyocardial support device should preferably be more compliant than themyocardium.

TABLE 1 Relative (uniaxial) Stiffness* Component (lbs/in) ReferenceMyocardium 3.8 to 5.0 Sideman & Beyar, “Simulation and Control of theCardiac System,” CRC Press, Inc., 1987, Chapter 5. Normal  0.1 to 25.0Lee et al., “Biaxial mechanical properties Human of human pericardiumand canine Pericardium comparisons, “Am. J. Physiol., 1987, 253:H75-H82.Latissimus 0.5 to 0.7 Reichenbach et al., “Passive Dorsi characteristicsof conditioned skeletal muscle for ventricular assistance,” ASAIO J.,1999 Jul.-Aug.; 45(4):344-9. Cardiac 0.8 to 1.7 Bench testing SupportDevice (i.e., the prior knit device) Bard Marlex  5.9 to 25.0 Benchtesting *Notes: Slope of stress versus strain curve (i.e. σ/ε) is ameasure of stiffness (lbs/in²). Incorporating the material thickness(t), a measure of relative stiffness is given by σ/ε (lbs/in). This is ameasure of load per inch width of material required to produce a givenstrain. Compliance is the inverse of stiffness. The uniaxial stiffnessfor various materials were derived from references listed for strains upto 30%.

In the Kass et al. article mentioned in the Background section, it wasspeculated that an artificial elastic “sock” could be used to replacethe latissimus dorsi muscle in adynamic cardiomyoplasty. This referenceseems to use the term elastic relative to compliance (not in it's trueengineering sense) and makes the comparison to replacing the latissimusdorsi with a device of similar compliance. The cardiac support device inTable I has compliance that is comparable or slightly less than thelatissimus dorsi muscle. However, the inventors have found that a highcompliance cardiac support device may have superior performance. Theadvantages of a high compliance cardiac support device (i.e., a devicehaving a compliance greater than the latissimus dorsi) are not disclosedby Kass et al.

Compliance of cardiac support devices can be measured in vitro todetermine either uniaxial directional compliance or 3-dimensional fulldevice volumetric compliance. The uniaxial directional compliance can bedetermined by taking samples of a selected device. These samples can bemounted on a standard hydraulically actuated tensile testing machinesuch as those supplied by MTS Systems Corporation or InstronCorporation. The compliance or stiffness characteristics of the deviceare obtained by measuring the load versus deflection of the sample. Thedevice and Bard Marlex stiffness provided in Table 1 were determinedusing this method for comparison purposes.

In use, the compliance of the cardiac support device is morerealistically based on 3-dimensional loading than uniaxial loading.Thus, an in vitro test was developed to examine full device compliance.For this test, the circumference of the base end of a sample cardiacsupport device (i.e., the prior knit device) was mounted to a supportplate to simulate the attachment of a device near the heart valve plane.A balloon bladder was placed in the volume defined by the device andfilled with saline to simulate the external heart ventricular volume. Tooffset the effect of the weight of fluid within the balloon, the mounteddevice and balloon are suspended in a tank of saline maintained atapproximately 37° C. The apex of the device is supported so when thefluid is added to the balloon, the device expands primarilycircumferentially, to better simulate the dilatation of a heart infailure. The internal volume of the balloon was monitored by recordingthe volume of fluid that was added incrementally. At each fluidincrement the pressures within the balloon are monitored using a Millarcatheter tip transducer and between the device and the external surfaceof the balloon using a “pillow” device and methods similar to thosedescribed by Tyberg et al. (“Static and dynamic operatingcharacteristics of a pericardial balloon,” Hamilton D R, Devries G,Tyberg J V, J Appl Physiol, April 2001;90(4):1481-1488). Both theinternal balloon pressure and pillow pressures track very closely,indicating very little resistance from the balloon.

Typical compliance curves obtained using this method with normalizedpercentage volume changes are shown in FIG. 10. FIG. 10 illustrates the3-dimensional or volumetric compliance curves for two cardiac supportdevice configurations. Curve A illustrates a lower compliance device,while curve B represents a higher compliance device.

Data indicates that a high compliance cardiac support device may bedesirable in many circumstances. As discussed above, the compliance of acardiac support device can vary over a given range of displacement, ordepending whether or not the device is subject to uniaxial or multiaxialloads.

As used herein, the term “high compliance cardiac support device” refersto a device having a compliance that is greater than that of a normalmyocardium, and more preferably greater than the compliance of thelatissimus dorsi muscle. In one embodiment, or characterization, thehigh compliance device of the invention can thus be described as havinga stiffness less than 3.8 lbs/in for uniaxial strains up to 30%. Asshown by the data in Table 1 and the discussions above, it may be morepreferable that the device has a compliance that is greater than anormal Latissimus Dorsi muscle, i.e., a stiffness less than 0.5 lbs/infor uniaxial strains up to 30%. Typically, when referring to a “highcompliance” device herein, the inventors are referring to a cardiacsupport device having a stiffness less than about 0.5 lbs/in whensubjected to a uniaxial load at strains up to 30%, more typicallybetween about 0.05 lbs/in and about 0.2 lbs/in. It will be appreciatedthat the foregoing description of data for strain up to 30% is intendedto be representative and not to suggest strains greater than 30% are notapplicable to the present invention.

Another way in which the compliance of the high compliance device can becharacterized is based on 3-dimensional volumetric compliance in termsof the percentage of volume increase (%) over applied pressure (mm Hg).Using this characterization in a representative example, the highcompliance device of the invention will have a compliance that allows atleast a 3% increase in volume for every 1 mm Hg increase in pressure(3%/mm Hg), more preferably between about 5%/mm Hg and about 15%/mm Hg.Actual volumes will depend upon the specific compliance. Again, it willbe appreciated the foregoing is a non-limiting example.

i. Material

The high compliance cardiac support device of the invention can befabricated using various materials and configurations to provide themechanical characteristics desired. In a preferred configuration, thedevice is constructed from a warp knitted fabric 18 of polyester fibers.Generally, the fabric 18 material is formed from intertwined fibers 20that are made up of a plurality of filaments 30, as shown in FIG. 9. Thecompliance of the material may be due to a variety of factors,including, but not limited to, the compliance of the individualfilaments 30 that make up the fibers 20 (see section b. Elasticity), therelative movement of the filaments 30 within a fiber 20, and/or therelative movement of the intertwined fibers 20 when subjected to load.Texturizing of the yarn can impact the compliance and elasticity of thefibers. Preferably, the fiber material and texturizing result in acompliant and elastic fiber such as a stretch polyester.

Compliance due to the relative movement (e.g., geometric deformation ofthe fabric openings) of the intertwined fibers 20 may be affected by themanner in which the fibers 20 are entwined. For example, a knit materialwill tend to be more compliant than a woven material because the loopsof the knit are capable of deforming (e.g., widening or lengthening) toaccommodate applied stress. In comparison, woven materials tend to haveless elongation unless elastomeric fibers are used. Knit material alsotends to recover well from deformation because the loops attempt toreturn to their original positions. The looped configuration of thefibers accommodates this recovery more readily than does the interwovenconfiguration found in woven materials. The ease and quickness withwhich elastic recovery takes place is also dependent on the fibercomposition. The fibers 20 of the jacket 10 material may be entwined asa knit (for example, a warp knit) or as a weave. Preferably, the fibers20 of the jacket 10 material are entwined as a knit.

ii. Manufacturing a High Compliance Device

The compliance of the cardiac support device can be due to theintertwining of the fabric fibers, or due to the compliance/elasticityof the fibers themselves, as discussed above. Additionally, thecompliance of the cardiac support device can be altered by the method ofprocessing the fabric.

For example, the compliance of the material of a cardiac support devicecan be increased by “shrinking” the material of the device, such thatthe device then includes more material within the same unit area and thefibers are closer together and more compressed, as compared to thedevice before the “shrinking” process. For example, shrinking can beaccomplished by heating the device. A memory condition can be introducedby a high temperature exposure or set temperature within the fibers,which modifies the “resting” state of the fibers (i.e., the state towhich they naturally return without the use of force). Exposure totemperatures below the set temperature can cause the fibers to respondby shrinking to the at rest condition. However, exposure to newtemperature conditions above the original set temperature whilesubjected to a load will create a new at rest configuration.Additionally, changing the fabric knit configuration, fiber texturizingor fiber material can further increase the compliance of the originaldevice.

Thus, in one embodiment, a high compliance device is manufactured byadding additional material to a fabric pattern of a lower compliancedevice (also referred to as the “original” device). Both patterns areshrunk to the same size, for example, using a heat set mandrel (i.e.,the same heat set mandrel is used for the “original device” and the“highly compliant” device). This method can easily increase thecompliance of the device 5 to 10 times (at low to moderate strains) overthe original device.

In manufacturing, the device is shaped to that of a healthy heart sothat the device not only uses its stored energy to reduce size but alsoto help the patient's heart restore shape. Both beneficial attributesare referred to as remodeling.

b. Elasticity

As used herein, the term “elastic” refers to the ability of the deformedmaterial to return to its initial state after a deforming load isremoved. A device that is highly elastic can undergo very largedeformations, but upon unloading returns to or close to its originalstate. With respect to a cardiac support device, elasticity may beimportant to the cardiac support device for maintaining an external loadon the heart as it reduces in size.

When a material is subjected to a deformation, the deformation is eitherplastic or elastic. If the deformation is plastic, it does not reboundwhen unloaded. The degree of elasticity for a given loading can becharacterized as the percentage of the deformation that rebounds uponunloading. Thus when unloaded, an entirely elastic material wouldrebound to its original state and characterized as 100% elastic at thatload. Whereas, a material that does not rebound at all from its deformedstate would be considered to have undergone an entirely plasticdeformation and would be considered 0% elastic at that load. In general,the amount of elastic recovery for the cardiac support device (in %) canbe calculated as 100%(d₁−d₂)/d₁, where d₁ is the initial deformation andd₂ is the deformation after unloading. The deformations d₁ and d₂ can bebased on any dimensional measure of length, area or volume as long asthe units are consistent.

Preferably the device 10 has an elastic recovery of at least about 50 %.However, it should be at least enough to allow the device to deformelastically up to the desired reduction in cardiac dimension targetedfor the chronic potential energy release. Thus, if the device isimplanted at 50% fabric strain and the desired heart size is calculatedto be at a point of 25% fabric strain, it would be preferable to have adevice capable of at least about 50% elastic recovery, more preferablyat least about 70% elastic recovery.

As with compliance, the elasticity of the material may be due to avariety of factors. The elasticity of the base material used tofabricate the device is one factor in determining the elastic recovery.For the cardiac support device, one suitable material is polyethyleneterephthallate (PET), more commonly known as polyester. Otherbiologically compatible materials could also be used to provide thedesirable amount of elasticity. In addition to the base material, theconfiguration and heat-induced memory are also important in determiningthe elasticity of the device. In one embodiment, a warp knitted fabricfabricated from continuous multi-filament set textured yarns is used.The fabric knit configuration contributes to the elastic performance ofthe device as well as its compliance. However, the process oftexturizing the yarn fibers 20 introduces a permanent crimp in the yarnthat is very important to the compliance and elastic performance of thefinal device.

The permanent crimp induced in the individual filaments 30 that make upthe yarn fibers 20 during texturizing provides a memory to the yarn. Thepermanent crimp can be deformed when loaded, but will have a tendency toreturn to the crimped configuration when unloaded (i.e. elasticallyrecover). The texturizing process generally involves heat anddeformations to form the permanent crimp. Stretching the fabric andheating to a higher temperature during the final processing of thedevice can remove some of the yarn crimp and provide a new memorycondition.

The preferred permanent yarn crimp for the original cardiac supportdevice is produced by set texturizing the yarns, then processing thefinal device by heat setting it with the device slightly stretched.Increased compliance and elasticity can be obtained using the samepolymer and fabric knit configuration, but with no final device heat setor by using other texturizing processes such as stretch textured yarns.As mentioned, in the preferred configuration elastic recovery is atleast 50%, but most preferable 70% to 100%.

Device elasticity can be determined from the compliance curves forloading and unloading a device. The in vitro 3D balloon compliance testdescribed in the previous section can be used to load and unload thedevice to determine the elastic rebound.

D. Benefits

The device 10 of the invention may provide some or all of the followingbenefits.

1 Reduction in Heart Dimension

The device 10 of the invention is a highly compliant and elastic devicethat is capable of mechanically reducing the heart size over time byusing the chronic potential energy release mechanism previouslydescribed. The reduced heart size is beneficial due to reduced wallstresses, which may, in turn, lead to improved cardiac function. Thebenefit of reducing heart size with a high compliance device can beillustrated by comparing a lower compliant device to a high compliancedevice.

FIG. 11 shows the compliance curves of two devices as implanted over theventricular portion of the heart. The lower compliance (A) and highercompliance (B) devices are both stretched to apply the same externalpressure (approximately 6 mm Hg) to the ventricular portion of the heartat the time of implant and initial heart volume (i.e. 0% heart volumeincrease). The devices plotted in FIG. 11 are the same devices as shownin FIG. 10. The zero points on the horizontal axes on the two Figuresare not the same. Therefore percent calculations between the Figureswill differ. In FIG. 10, the zero point is an “at rest” value for thedevice. In FIG. 11, the zero point is after implantation. The highercompliance of the “B” device was obtained by adding more material to thedevice so that less yarn crimp was removed during heat setting. Thelower compliance of the “A” device is indicated by a steeper line. Thecompliance of either device at any point of either curve can beexpressed as 1/slope of the curve at that point. Thus, at implantation,the compliance of the higher compliance device (B) is 5.5%/mm Hgcompared to about 2%/mm Hg for the less compliant (A) device. If theoperating range of the device is assumed to be below a 20% volumeincrease, the compliance range is between about 3%/mm Hg and 20%/mm Hgfor the highly compliant (B) device and between about 1%/mm Hg and3.5%/mm Hg for the less compliant (A) device. In this example, thehigher compliance device is approximately 3 to 6 times more compliantdepending on the given condition within the operating range. As usedherein, a “high compliance” device refers to a device having acompliance between about 3%/mm Hg and about 20%/mm Hg , or greater. A“low compliance” device refers to a device having a compliance betweenabout 1%/mm Hg and about 3%/mm Hg, or lower. In FIG. 11, the devices cantheoretically apply an external pressure to the heart until the heartvolume decreases to the point where the compliance curve crosses 0 mmHg. These curves are based on loading, not unloading. Therefore, as theheart volume decreases, these curves assume that both devices have 100%elastic recovery. In general, even though the two devices do not have100% elastic recovery, for comparison purposes the higher compliancedevice will have better elastic recovery than the low compliance device.Whereas a “low compliance” device may have the potential to reduce thesize of a heart between about 10% to about 20% in volume, a highcompliance device can continue to apply an external load to the heart toachieve up to between about a 50% to about a 60% volume reduction.

Depending on the heart shape change that is assumed (i.e., cylindricalor spherical), the volume decrease for the “low compliance” devicecorresponds to a decrease in diameter between about 5% to about 10%. Thevolume decrease for the “high compliance” device similarly correspondsto a decrease in diameter between about 15% to about 30%.

A “low compliance device” corresponding to the lower compliance device(A) has been implanted in human clinical trails with follow-up out to 12months post-implant (i.e., the prior knit device). FIG. 13 shows theaverage change in left ventricular end diastolic diameter (LVEDD) for 17patients receiving the lower compliance device. At implant the heartswere fit to provide acute support that resulted in a 5.2% reduction inLVEDD. After 3 months post-surgery, the LVEDD decrease another 4.8% onaverage. This additional chronic reduction in LVEDD corresponds closelywith the 5% to 10% diameter reduction of the external ventricular sizepredicted by the device compliance curve shown in FIG. 11 and thechronic potential energy mechanism. In fact, the amount of elasticrecovery for the low compliance device has been measured in vitro to beapproximately 70% to 80%, depending on the loads applied.

A typical loading and unloading curve for a lower compliance device isshown in FIG. 14. The elastic recovery calculated from FIG. 14 isapproximately 80%. This was calculated based on the percentage of thevolume change from initial to fully deformed that was recovered (i.e.fully deformed to new unloaded resting state volume). If the predicteddiameter reduction range of 5% to 8% is reduced to account for less than100% elastic recovery, the expected decrease in diameter would bebetween 3.5% and 6.5%. The actual clinical result is nearly in themiddle of this predicted range.

The potential reduction in heart size attributable from the chronicpotential energy release mechanism can also be examined based on theenergy that is stored in the device relative to the device compliance.Elastic potential energy stored in a spring is equal to the amount ofwork energy (U) used to compress it if no frictional or other losses areassumed. Thus, the work energy can be determined as follows:

-   -   U=work energy=Fx/2    -   Where:        -   F=applied force=Kx        -   K=stiffness=1/compliance        -   X=displacement

FIG. 12 illustrates the potential energy stored in both a low compliance(A) and high compliance (B) device at implant. Both energy curves assumethat the device is implanted on the same size heart (i.e. externaldiameter of 8.4 cm.) with the same externally applied pressure of 6 mmHg when implanted. The high compliant device has nearly 4 times moreenergy (508 mJ versus 131 mJ) at the time of implant. Although theamount of energy stored is due to the compliance, the amount availablefor release to reduce the heart size is based on the elastic recovery ofthe device. For example, if the device has 80%,elastic recovery, then80% of the energy is available to drive the heart smaller, while 20% ofthe energy is lost to permanent deformation of the material.

2. Eliminate Surgical Fit

As discussed previously, in one embodiment, the cardiac support deviceis adjusted at the time of implantation to provided the desired fit. Theadjustment of the device allows it to be used on dilated hearts having alarge range of shapes and sizes. To accommodate such variability indilated hearts, the device is manufactured in many sizes. However, manydozens of sizes would be necessary to provide a sufficient selection forproviding the appropriate fit across all the range of heart shapes andsizes.

One advantage of a “high compliance” jacket is that each jacket canadapt to a large shape/size range, yet still provide the appropriatefit. Since compliance is defined as the deformation for a given load, ahigh compliance device will result in a large change in deformation witha relatively small change in load. Thus, a target load or fit range cantheoretically be accommodated by a larger displacement range with a highcompliance device than for a low compliance device.

For example, two device compliance curves, low compliance (A) and highcompliance (B) are shown in FIG. 10. If the chosen target fit load (i.e.pressure applied to the epicardial surface of the heart) is betweenabout 5 mm Hg to about 7 mm Hg of pressure, the dimensional stretchrange from the device resting state can be determined from the curvesfor each device. Device “A” can be stretched anywhere between a 17% to23% (6% range) increase in volume from it's starting volume and willapply a 5 mm Hg to 7 mm Hg pressure. However, the higher compliancedevice “B” can be stretched over a larger range of 129% to 156% (27%range) from it's starting volume for the same resulting load. Nowsuppose it is desirable to manufacture devices that will apply a 5 mm Hgto 7 mm Hg load for heart sizes from 645 ml to 780 ml of externalventricular volume. The high compliance device (B) would require onlyone size to cover the range of heart sizes selected, but it would take 4sizes of the low compliance device (A) to accommodate the heart sizerange. This example is illustrated in Table 2.

TABLE 2 No. Device Size (@ rest) Min. Volume Max. Volume Sizes Device(ml) (ml) (ml) Required B 500 645 780 1 A 550 644 677 4 578 676 711 607710 746 637 745 783

Eliminating surgical fit based on a high compliance device may bebeneficial for several reasons. First of all, although the surgicalprocedure for implanting a low compliance device is relatively simplecompared to other cardiac surgeries, eliminating the fitting processwould further simplify the surgery. The surgical fitting step is one ofthe most time-consuming steps of the surgical implant process.Eliminating this step could shorten the overall surgical time. Thiswould result in the patient undergoing anesthesia for a shorter periodof time, reducing the risks due to anesthesia dose complications. Inaddition, the reduced surgical time could reduce the overall surgicalcosts due to a reduction in the time spent in the operating room.

Another benefit of eliminating the surgical fit is increasedconsistency. Surgically adjusting and fitting each device tends tointroduce variability between patients by any given surgeon. Inaddition, there is variability between different surgeons and hospitalsthat can only be reduced by rigorous implant training. Thus, eliminatingthe surgical fit procedure may reduce the variability and the trainingrequirements.

Eliminating the surgical fit can also make implanting the device morecompatible with minimally invasive surgical approaches. Typically, thesurgical fit step for implant of the original device requires access tothe anterior portion of the heart. This is most commonly accomplishedusing a full median sternotomy surgical approach. If a high compliancedevice can allow the appropriate fit to be obtained through device sizeselection rather than surgical customization of fit, the surgery may bepossible through a smaller incision than a full median sternotomy. Itmay even be possible to implant the device through small portalincisions.

Minimally invasive surgical incisions can have numerous benefits,including reduced pain, less cosmetic scaring, faster hospital releaseand faster return to physical activities. The reduced hospital stay fromminimally invasive surgery can also help to reduce overall surgicalcosts and make the surgery more accepted and routine in the medicalcommunity.

In one embodiment, the cardiac support device is 3-dimensional shapethat is constructed from a flat fabric mesh. To form a 3-dimensionalshape from a flat fabric, the device typically includes sewn seams wherethe device material is a little denser and thicker. Unfortunately, theseams can result in a greater tissue response and an increased potentialfor adhesion between the device and adjacent tissues in the chest, otherthan the heart.

Thus, in another embodiment, the device is manufactured using advancedfabrication methods that eliminate the manufactured seams. However, evenif the pre-fabricated seams can be eliminated, the surgical fitting mayresult in a seam that is even thicker and denser than those producedduring device manufacture. Consequently, if the surgical seam fromfitting is eliminated, for example, by using a high compliance device,and the manufactured seams are eliminated, the potential for adhesionsto adjacent tissues would be reduced. This could be important,particularly when future surgeries require access to the chest cavity.Adhesions make surgical access much more difficult.

3. More Volume Overloading Tolerant

Another advantage of a high compliance device is the ability of thedevice to expand and not over restrain the heart in the case of volumeoverloading. For example, excessive fluid intake can impact the volumeof the heart. A high compliance device may benefit the patient byhelping to support the increased volume loading, without overlyrestricting the heart as might occur with the cardiac condition known asconstriction. Although dilatation constraint is a potentially importantmechanism of the cardiac support device, higher compliance may provideadequate support and resistance to dilatation without overly restrictingthe patient's normal variations in fluid intake.

Having disclosed the invention in a preferred embodiment modificationsand equivalents will become apparent to those skilled in the art. It isintended such modifications and equivalents shall be included within thescope of the appended claims. For example, while the invention isdescribed covering the ventricles, the invention can cover one or bothof the atria only or in combination with ventricle coverage. Also, thedevice can be provided with circumferential fibers which have a maximumstretch (or no stretch) at a volume representing a maximum volume forend diastole at time of placement. Such a modification provides acuteprevention of diastolic expansion beyond a maximum. Use of multiple setsof fibers are described in Haindl international patent application PCTWO 98/58598 published Dec. 30, 1998.

1. A device for treating diseases of a heart, said device comprising: abiocompatible material sized to surround an external surface of saidheart; said material having an elasticity and a compliance reflecting atendency to return to a rest state and an ability to deform understrain, respectively; said elasticity and compliance selected for saidmaterial to be placed surrounding said external surface with a tightnessselected for energy to be stored in said material to assist chronicremodeling of said heart and to avoid significant acute resistance todiastolic filling of said heart, wherein compliance is the inverse ofstiffness, and wherein the biocompatible material has a stiffness ofless than about 3.8 lbs/in when subjected to a uniaxial load at a strainof less than 30%.
 2. The device according to claim 1 wherein thematerial has a compliance greater than a compliance of a normalmyocardium, wherein said biocompatible material conforms to an externalsurface of the heart and is sized to provide resistance tocircumferential expansion of the heart without impeding systoliccontraction.
 3. The device according to claim 1 wherein the device has acompliance greater than a compliance of a normal latissimus dorsi,wherein said biocompatible material conforms to an external surface ofthe heart and is sized to provide resistance to circumferentialexpansion of the heart without impeding systolic contraction.
 4. Thedevice according to claim 3, wherein the material has a stiffness ofless than about 0.5 lbs/in when subjected to a uniaxial load at a strainof less than 30%.
 5. The device according to claim 1, wherein thebiocompatible material has a stiffness of less than 0.5 lbs/in whensubjected to a uniaxial load at a strain of less than 30%.
 6. The deviceaccording to claim 1, wherein the biocompatible material has a stiffnessbetween about 0.05 lbs/in and about 0.2 lbs/in when subjected to auniaxial load at a strain of less than 30%.
 7. The device according toclaim 1, wherein the material has a compliance that allows at least a 3%increase in volume for every 1 mm Hg increase in applied pressure. 8.The device according to claim 1, wherein the material has a compliancebetween about 5%/mm Hg and about 15%/mm Hg.
 9. The device according toclaim 1, wherein the material is sized to be smaller than the externalsurface of the heart to which it is applied, wherein the material isconfigured to exert a pressure on the external surface of the heart thatis no greater than an end diastolic pressure of a right ventricle of theheart.
 10. The device according to claim 9, wherein the material definesa resting volume and is stretched to provide a stretched volume that isat least 20% greater than the resting volume.
 11. The device accordingto claim 1, wherein the material is sized to be larger than the externalsurface of the heart to which it is applied and adapted to be sized byadjustment during implantation.
 12. The device according to claim 1,wherein the material is capable of an elastic recovery of at least about50%.
 13. The device according to claim 1, wherein the material iscapable of an elastic recovery of at least about 70%.
 14. The deviceaccording to claim 1, wherein said material is configured to cover aventricular portion of the heart.
 15. The device according to claim 1,wherein said material is configured as a jacket having an upper and alower end, wherein said jacket is open at said upper and defines aninternal volume between said upper and lower end capable of receivingthe heart.
 16. The device according to claim 15, wherein said materialis closed at said lower end.
 17. The device according to claim 15,wherein said material is open at said lower end.
 18. The deviceaccording to claim 1, wherein said material is configured to cover atleast one ventricle of said heart.
 19. The device according to claim 18,wherein said material is configured to cover a left ventricle of saidheart.
 20. The device according to claim 18, wherein said material isconfigured to cover a right ventricle of said heart.
 21. The deviceaccording to claim 18, wherein said material is configured to cover aright and a left ventricle of said heart.
 22. The device according toclaim 1, wherein said material is configured as a patch.
 23. The deviceaccording to claim 1, wherein said material comprises intertwinedfibers.
 24. The device according to claim 23, wherein said materialcomprises a knit.
 25. The device according to claim 23, wherein saidmaterial comprises a weave.
 26. The device according to claim 23,wherein said fibers are crimped.
 27. The device according to claim 26,wherein said crimped fibers are produced by stretch texturizing.
 28. Thedevice according to claim 27 wherein said fibers consist of a stretchtextured polyester or PET or other polymeric materials.
 29. The deviceaccording to claim 1, wherein said material comprises polyethyleneterephthalate (PET) or other polymeric or biological materials with saidproperties.
 30. The device according to claim 1, wherein said materialis configured to apply a pressure to the external surface of the heartat end diastole of less than 10 mm Hg.
 31. The device according to claim1, wherein said material is configured to apply a pressure to theexternal surface of the heart at end diastole of between about 5 mm Hgand about 7 mm Hg.
 32. A method for treating diseases of a heart, saidmethod comprising: surgically accessing the heart; applying a cardiacsupport device to an external surface of the heart where said cardiacsupport device includes: a biocompatible material sized to surround anexternal surface of said heart; said material having an elasticity and acompliance reflecting a tendency to return to a rest state and anability to deform under strain, respectively; said elasticity andcompliance selected for said material to be placed surrounding saidexternal surface with a tightness selected for energy to be stored insaid material to assist chronic remodeling of said heart and to avoidsignificant acute resistance to diastolic filling of said heart, whereincompliance is the inverse of stiffness, and wherein the biocompatiblematerial has a stiffness of less than about 3.8 lbs/in when subjected toa uniaxial load at a strain of less than 30%.
 33. The method accordingto claim 32 wherein said remodeling includes a reduction in a volume ofsaid heart.
 34. The method according to claim 32 wherein said remodelingincludes an altering of a shape of said heart.
 35. The method accordingto claim 32 wherein said material has a compliance greater than thecompliance of a normal latissimus dorsi, wherein said biocompatiblematerial conforms to an external surface of the heart and is sized toprovide resistance to circumferential expansion of the heart withoutimpeding systolic contraction.
 36. The method according to claim 32,wherein the material is configured to resist dilatation of the heart.37. The method according to claim 32, wherein the material is sized tobe larger than the external surface of the heart to which it is appliedduring diastole, said method further comprising a step of adjusting asize of said material to conform to the external surface of the heart.38. The method according to claim 37, wherein the material is adjustedto exert a pressure on the external surface of the heart of less thanabout 10 mm Hg.
 39. The method according to claim 37, wherein thematerial is adjusted to exert a pressure on the external surface of theheart of less than about 5 mm Hg.
 40. The method according to claim 37,wherein the material is adjusted to exert a pressure on the externalsurface of the heart of less than about 2 mm Hg.
 41. The methodaccording to claim 32, wherein said material is configured to acutelysupport the external surface of the heart without custom fitting. 42.The method according to claim 41, wherein said material is sized to besmaller than the external surface of the heart to which it is appliedduring diastole.
 43. The method according to claim 42, wherein saidmaterial exerts a pressure on the external surface of the heart that isless than an end diastolic pressure of a right ventricle of the heart.44. The method according to claim 42, wherein said material isconfigured to impose between about a 5% to about a 10% reduction inmaximum diastolic volume of the heart upon implantation.
 45. The methodaccording to claim 42, wherein said material is configured to exert apressure on the external surface of the heart at end diastole betweenabout 2 mm Hg and about 20 mm Hg.
 46. The method according to claim 42,wherein said material is configured to exert a pressure on the externalsurface of the heart at end diastole between about 5 mm Hg and about 15mm Hg.
 47. The method according to claim 42, wherein said material isconfigured to exert a pressure on the external surface of the heart atend diastole between about 5 mm Hg and about 10 mm Hg.
 48. The methodaccording to claim 32, wherein said material is configured to enhance areduction in a size of the heart.
 49. The method according to claim 48,wherein said material is sized to be smaller than the external surfaceof the heart to which it is applied during diastole.
 50. The methodaccording to claim 49, wherein the material defines a resting volume,said method further comprising a step of stretching the material toprovide a stretched volume that is at least 20% greater than the restingvolume.
 51. The method according to claim 50, wherein said materialexerts a pressure on the external surface of the heart that is less thanan end diastolic pressure of a right ventricle of the heart.
 52. Themethod according to claim 50, wherein said step of stretching thematerial provides the material with potential energy to mechanicallyreduce a size of the heart.